This laser apparatus, as well as the method of material treatment on which it is based, are particularly useful in forming curved cuts within a transparent material. Curved cuts within a transparent material are generated, for example, in laser-surgical methods, in particular in eye surgery. This involves focusing the treatment laser radiation within the tissue, i.e. beneath the tissue surface, so as to form optical breakthroughs in the tissue.
In the tissue, several processes initiated by the treatment laser radiation then occur in a time sequence. If the power density of the radiation exceeds a threshold value, an optical breakthrough will result, generating a plasma bubble in the material. After the optical breakthrough has formed, said plasma bubble grows due to expanding gases. Subsequently, the gas generated in the plasma bubble is absorbed by the surrounding material, and the bubble disappears again. However, this process takes very much longer than the forming of the bubble itself. If a plasma is generated at a material boundary, which may quite well be located within a material structure as well, material will be removed from said boundary. This is then referred to as photo ablation. In connection with a plasma bubble which separates material layers that were previously connected, one usually speaks of photo disruption. For the sake of simplicity, all such processes are summarized here by the term optical breakthrough, i.e. said term includes not only the actual optical breakthrough, but also the effects resulting therefrom in the material.
For a high accuracy of a laser surgery method, it is indispensable to guarantee high localization of the effect of the laser beams and to avoid collateral damage to adjacent tissue as far as possible. It is, therefore, common in the prior art to apply the laser radiation in a pulsed form, so that the threshold value for the power density of the laser radiation required to cause an optical breakthrough is exceeded only during the individual pulses. In this regard, U.S. Pat. No. 5,984,916 clearly shows that the spatial extension of the optical breakthrough (in this case, of the generated interaction) strongly depends on the pulse duration. Therefore, high focussing of the laser beam in combination with very short pulses allows to place the optical breakthrough in a material with great point accuracy.
The use of pulsed laser radiation has recently become established practice particularly for laser-surgical correction of visual deficiencies in ophthalmology. Visual deficiencies of the eye often result from the fact that the refractive properties of the cornea and of the lens do not cause optimal focusing on the retina.
U.S. Pat. No. 5,984,916 mentioned above, as well as U.S. Pat. No. 6,110,166, describe methods of producing cuts by means of suitable generation of optical breakthroughs, so that, ultimately, the refractive properties of the cornea are selectively influenced. A multitude of optical breakthroughs are joined such that a lens-shaped partial volume is isolated within the cornea of the eye. The lens-shaped partial volume which is separated from the remaining corneal tissue is then removed from the cornea through a laterally opening cut. The shape of the partial volume is selected such that, after removal, the shape and, thus, the refractive properties of the cornea are changed so as to have the desired correction of the visual deficiency. The cuts required here are curved, which makes a three-dimensional adjustment of the focus necessary. Therefore, a two-dimensional deflection of the laser radiation is combined with simultaneous adjustment of the focus in a third spatial direction. This is summarized herein by the term “deflection”.
When forming a cut by joining optical breakthroughs in the material, an optical breakthrough is generated several times faster than the time it takes until a plasma generated therefrom is absorbed by the tissue again. It is known from the publication by A. Heisterkamp, et al., in: Der Ophthalmologe, 2001, 98:623-628, that a plasma bubble grows after an optical breakthrough has been generated in the cornea of the eye at the focal point where the optical breakthrough was generated, which plasma bubble reaches a maximum size after a few ns and then almost completely collapses again. This leaves only small residual bubbles. Said publication states that joining of still growing plasma bubbles will reduce the quality of the cut. Therefore, it suggests a method of the above-mentioned type, wherein individual plasma bubbles are not generated directly next to each other. Instead, a gap is left between two sequentially generated optical breakthroughs, which breakthroughs are generated along a spiral-shaped path. The gap is filled, in a second run, along the spiral with optical breakthroughs and with plasma bubbles resulting therefrom. This is intended to prevent adjacent plasma bubbles from being connected with each other and to improve the quality of the cut.
However, it is generally required to control the distance between two subsequent plasma bubbles along the path line as precisely as possible. In the case of a constant repetition rate, this may be principally effected by adapting the feed speed, i.e. the speed of deflection. In the case of the spiral, this would mean that the laser beam passes along an inner spiral path portion at a much higher speed (i.e. a higher angular frequency) than along an outer path portion. This is a suitable method as long as the maximum frequency of deflection of the scanner system used permits a sufficient feed speed. For the frequency of deflection fs of the scanner used for lateral deflection of the laser beam, the simple relationship fs=(fL*s)/(2π*r) holds. In this relationship, fL is the repetition rate of the pulses in the pulsed laser beam and s is the geometrical distance, measured along the path line, between two plasma bubbles to be generated sequentially along an at least partially circular path line having a radius r. If the maximum frequency of deflection of common galvanometer scanners, which can follow the control signal with good precision up to frequencies of ca. 300 Hz in a non-resonant manner, is assumed for an estimation, this results in a maximum pulse frequency of about 4 kHz for s=10 μm and r=20 μm. With limitations concerning the angles of deflection, even higher pulse frequencies might possibly be put to reasonable use as well. However, this increases positional errors, thus setting practical limits to such procedure. These considerations show that, for presently common scanner systems, it is required to limit the pulse frequency of the laser radiation to a maximum of 10 kHz for generation of desired spiral paths.
As an alternative approach, it would be theoretically conceivable to make the pulse frequency of the laser radiation variable; however, there are certain limitations to such procedure when using laser systems having passively mode-synchronized oscillators. Therefore, for medical applications, the fs laser systems common today only provide laser radiation having a fixed pulse frequency. This leads to technical solutions which have fixed pulse frequencies of the laser radiation in the region of a few kHz. The process speed for generating the cuts is, thus, adapted to those regions of the path which place the highest demands on deflection.
Generating the cuts as quickly as possible is desirable not only for convenience or in order to save time; bearing in mind that movements of the eye inevitably occur during ophthalmological operations, quick generation of cuts also contributes to the optical quality of the result thus achieved and avoids the requirement to track eye movements.